The present invention relates to the magnetic resonance art. It finds particular application in conjunction with self shielded gradient coil assemblies for magnetic resonance imaging apparatus and will be described with particular reference thereto. However, it will be appreciated that the present invention will also find application in conjunction with magnetic resonance spectroscopy systems and other applications which require RF fields and gradient magnetic fields.
Heretofore, magnetic resonance imagers have included a superconducting magnet which generated a temporally constant primary magnetic field. The superconducting magnet was encased in a toroidal vacuum vessel or dewar that included a central cylindrical bore through which the primary magnetic field was generated. As generated, the primary magnetic field commonly had limited non-uniformities. Active or passive shims were mounted to the inside surface of the cylindrical bore to correct the non-uniformities and render the primary magnetic field substantially uniform within the bore. Although active shims in the form of resistive coils were sometimes used, passive shims in the form of small pieces of steel about an inch square and 0.010 inches thick were more common.
There are several important subsystems of the MRI system that are placed within the cylindrical bore of the primary magnet. Usually the structure closest to the bore is a shimset. The shimset may consist of coils carrying a specified amount of current or ferromagnetic material placed so as to reduce the non-uniformity of the primary magnetic field. The next inward structure is a set of coils for generating the x, y, and z-gradient fields used for MRI. Within the gradient coil structure is the RF coil and the object to be imaged. One alternative to this arrangement has been to place superconducting shim coils within the cold region of the magnet. This placement of shimset serves to separate the shims as far as possible from the imaging volume which is desirable when removing low order field inhomogeneity.
One of the difficulties was that the gradient magnetic fields generated for the MRI also induced eddy currents in the magnet structures. These eddy currents in turn produced their own magnetic fields thus interfering with the imaging process. The induced eddy currents can be compensated for with pre-emphasis only to the extent that the pre-emphasis perfectly mimic the gradient fields. Since there are some very cold conductive surfaces within the crystal to minimize helium boil-off, eddy currents could persist for seconds. Separation between the gradient coils and the magnet structure reduced the contribution of eddy currents and allowed plenty of space for a shimset.
More recently, self shielded gradients have been developed so as to eliminate eddy currents. A self shielded gradient consists of a set of primary x, y, and z coils and an additional set of x, y, and z-coils at a larger radius. The geometry of the coils is chosen such that when the coils are excited in series, they have substantially no residual gradient field outside of the outer coils. Within the inner radius, the combination of coils produces a substantially linear gradient. A larger separation between the inner and outer coils of a shielded gradient results in more efficient gradient field generation.
The cost of a large bore superconducting magnet is generally more than a smaller magnet. Thus, reduced system cost as well dictates minimum spacing between the gradient shield coil and the magnet bore.
The present invention provides a new and improved self shielded gradient coil design which overcomes the above-referenced problems and others.